1. Field of the Invention
This invention relates to a gamma camera sensitivity correcting device and method that corrects the sensitivity of a gamma camera.
2. Discussion of Background
When a gamma camera is used, a subject is dosed with a radio-isotope (hereinbelow abbreviated to RI), the amount of this RI that is selectively accumulated in a specific organ or disease site is detected, and imaged as an accumulation image.
The main parts of a gamma camera are shown in FIG. 1. Reference numeral designates a detector that detects radiation (gamma rays) from the RI. It consists of a scintillator 17, light guide 18, photoelectron multiplier tubes 19, and pre-amplifiers 20. Gamma rays cannot be refracted and focussed by a lens as in an optical camera, so that a collimator 16 is arranged ahead of detector 15. Collimator 16 is built around a lead sheet provided with a large number of parallel holes. By this means an image of the same size as the subject is created at the surface of scintillator 17.
In scintillator 17, photoelectron multiplier tubes (PMTs) 19 are arranged in a hexagonally close-packed fashion on the other side of light-guide 18, which is constructed of a transparent substance (lucite or the like). A large number (19, 37 or 61) of PMTs 19, of 2-inch or 3-inch diameter, are used, arranged in a hexagonal pattern. When gamma rays are incident on scintillator 17, their energy is absorbed, causing fluorescence at the point of absorption. This light is input via light guide 18 to PMTs 19, where it is subjected to a photoelectric conversion, so that output pulses proportional to the incoming light are output from each PMT 19. If light is emitted directly below a given PMT, the largest output will be obtained from that PMT; or, alternatively, if light is emitted at a point between three PMTs, equal output will obtained from those three PMTs. Thus the point of luminescence, i.e., the position of gamma ray incidence, can be found from the output pulses of PMTs 19. Using the values of X and Y of an orthogonal coordinate system whose origin is at the center of scintillator 17, adder 21 multiplies the outputs of the respective PMTs 19 by coefficients (called "weights") that depend on the positional coordinates of the respective PMT and adds them, to obtain signals for the various directions X.sup.+, X.sup.-, Y.sup.+, and Y.sup.-.
In addition to this, if the outputs of all the PMTs 19 are multiplied by a constant coefficient and then added, the resultant signal will be proportional to the total luminescence of the scintillator, i.e., to the energy of the gamma ray. This is therefore termed the energy signal or Z signal. This Z signal is amplified by amplifier 22, then input to pulse wave height analyzer 25, so that only gamma rays of the required preset energy are selected. The X.sup.+, X.sup.-, Y.sup.+, and Y.sup.- position signals and the Z signal are each amplified and then the signals of the selected gamma rays only are input to a waveform expanding circuit 23, from the output of which signals: ##EQU1## are obtained. The reason for dividing by Z is to make the magnitude of the image independent of the gamma ray. These X and Y signals are applied to the deflecting input of a cathode ray oscilloscope (CRT). The output of pulse wave height analyzer 25 is applied to a spot generating circuit 26, so that each incoming gamma ray is displayed as a luminous point on CRT 27. A 2-dimensional image can be obtained by recording and accumulating these luminous points on polaroid film or X-ray film by means of a lens. It is also possible to record and accumulate the frequency with which gamma rays arrive in an IC memory corresponding to X and Y and to obtain a functional diagnostic image by carrying out data processing on the data stored in this IC. FIG. 2 shows a case where such a functional diagnostic image is obtained.
Position signals X and Y are input through A/D converters 2 and 3 to correction table 5 and linear correction part 6. From correction table 5, correction values ai(X,Y), bi(X,Y), ci(X,Y), and di(X,Y); (where i=1, 2, . . . ) corresponding to the X and Y signals are output, and these are then input to linear correction part 6. This linear correction part 6 finds X' and Y' by calculating EQU X'=a.sub.1 X+b.sub.1 Y+c.sub.1 XY+d.sub.1 EQU Y'=a.sub.2 X+b.sub.2 Y+c.sub.2 XY+d.sub.2
The Z signal is also input to correction part 7, where it is corrected by performing the calculation EQU Z'=e(X,Y) . Z
This corrected data Z' is input to window circuit 9, constituting the next stage, and, if this input Z' is of a prescribed energy value, it causes the pixel value of the corresponding address of image data memory 10 to be incremented by +1. The memory content of image data memory 10 is input to main CPU (central processing unit) 8 through line 11 and the values of Z, X, and Y are input to main CPU 8 through lines 12 and 13, where calculation and rewriting of the correction coefficients are performed. Main CPU 8 performs a sensitivity correction on the contents of image data memory 10 in accordance with correction coefficients f(i,j). In collection modes such as single photon emission CT (hereinbelow termed "ECT") where strict uniformity etc. is required, main CPU 8 performs sensitivity correction in accordance with: EQU N'(i,j)=f(i,j) X N(i,j)
where (i,j) are the memory addresses corresponding to (X,Y).
Since there is a timewise variation of the output wave height of photoelectron multiplier tube 19 due to optical coupling of scintillator 17 and light guide 18, or to variation of the characteristic of photoelectron multiplier 19 etc., it is necessary to re-write the correction data of Z correcting part 7 periodically, to match these changed circumstances.
Furthermore, being made of a metal, such as lead, of high gamma ray screening ability, collimator 16 shows differences in the proportion of gamma rays that it lets through, depending on the precision with which it was machined and on the location (X,Y). Consequently, in the case of ECT, there were the problems that ring-shaped artifacts were formed in the reconstructed image due to statistical scatter in the uniformity of the sensitivity of the gamma camera, or that artifacts were produced with lapse of time (however, such artifacts are not produced if the correction data are re-written to match time-wise variation, as mentioned above). This correction of uniformity of sensitivity for the ECT collection purposes was carried out by creating a correction table (matrix) from the uniformity data of detector 15 (scintillator+light guide+photoelectron multiplier)+electrical system+collimator 16, and implementing a sensitivity correction as (collimator+detector) by multiplying the data in image data memory 10 by these coefficients. In this case, a change in the sensitivity of the whole system is seen due to the time-wise variation of the sensitivity of detector 15. Conventionally, to obtain correction data, the following steps (1) to (3) are necessary:
(1) find the correction coefficient of the Z correction part 7 by removing the collimator; PA1 (2) find the correction coefficient of linear interpolation part 6 by mounting a phantom for X-Y correction (linear correction) on the detector; PA1 (3) find the uniformity correction matrix by collecting uniformity data with (collimator+detector) using a uniform linear source, by mounting collimator 16 on detector 15 after (1) and (2) above.
However, the data collection required for sensitivity correction of the detector with collimator fitted (step (3) above) usually takes 2 to 3 hours, and it is difficult to perform this data collection frequently, since there are a number of different types of collimators.